Radially expandable endoprostheses are artificial devices adapted to be implanted or deployed in an anatomical lumen. An “anatomical lumen” refers to a cavity, duct, of a tubular organ such as a blood vessel, urinary tract, and bile duct. Stents are examples of endoprostheses that are generally cylindrical in shape and function to hold open and sometimes expand a segment of an anatomical lumen (one example of a stent is found in U.S. Pat. No. 6,066,167 to Lau et al). Stents are often used in the treatment of atherosclerotic stenosis in blood vessels. “Stenosis” refers to a narrowing or constriction of the diameter of a bodily passage or orifice. In such treatments, stents reinforce the walls of the blood vessel and prevent restenosis following angioplasty in the vascular system. “Restenosis” refers to the reoccurrence of stenosis in a blood vessel or heart valve after it has been treated (as by balloon angioplasty, stenting, or valvuloplasty) with apparent success.
The treatment of a diseased site or lesion with a stent involves both delivery and deployment of the stent. “Delivery” refers to introducing and transporting the stent through an anatomical lumen to a desired treatment site, such as a lesion. “Deployment” corresponds to expansion of the stent within the lumen at the treatment region. Delivery and deployment of a stent are accomplished by positioning the stent about one end of a catheter, inserting the end of the catheter through the skin into an anatomical lumen, advancing the catheter in the anatomical lumen to a desired treatment location, expanding the stent at the treatment location, and removing the catheter from the lumen.
A self-expanding stent is capable of expanding from a compressed or collapsed state to a radially expanded state. A delivery device which retains the stent in its compressed state is used to deliver the stent to a treatment site through vessels in the body. After the stent is positioned at the treatment site, the delivery device is actuated to release the stent which allows the stent to self-expand within the body vessel. The delivery device is then detached from the stent and removed from the patient. The stent remains in the vessel at the treatment site as an implant.
The stent must be able to satisfy a number of basic, functional requirements. The stent must be capable of withstanding the structural loads, for example, radial compressive forces, imposed on the stent as it supports the walls of a vessel after deployment. Therefore, a stent must possess adequate radial strength. After deployment, the stent must adequately maintain its size and shape throughout its service life despite the various forces that may come to bear on it. In particular, the stent must adequately maintain a vessel at a prescribed diameter for a desired treatment time despite these forces. The treatment time may correspond to the time required for the vessel walls to remodel, after which the stent is no longer necessary for the vessel to maintain a desired diameter.
Radial strength, which is the ability of a stent to resist radial compressive forces, relates to a stent's radial yield strength around a circumferential direction of the stent. A stent's “radial yield strength” or “radial strength” (for purposes of this application) may be understood as the compressive loading, which if exceeded, creates a yield stress condition resulting in the stent diameter not returning to its unloaded diameter, i.e., there is irrecoverable deformation of the stent. When the radial yield strength is exceeded the stent is expected to yield more severely and only a minimal force is required to cause major deformation.
Even before the radial yield strength is exceeded there may be permanent deformation in the stent following a radial compressive load, but this degree of permanent deformation somewhere in the stent is not severe enough to have a significant effect on the stent's overall ability to radially support a vessel. Therefore, in some cases the art may view “radial yield strength” as the maximum radial loading, beyond which the scaffold stiffness changes dramatically. “Radial yield strength” units are sometimes force-divided-by-length, which is an expression of radial yield strength on a per-unit-length basis. Thus, for a radial yield strength per unit length, e.g., F N/mm, the radial load which, if it exceeds this value, would result in significant change in stiffness for a stent having two different lengths, L1 and L2, would therefore be the product F*L1 and F*L2, respectively. The value F, however, is the same in both cases, so that a convenient expression can be used to appreciate the radial yield strength independent of the length of the stent. Typically, the radial force that identifies the point where stiffness is lost does not change much on a per-unit-length basis when the stent length changes.
A radial “stiffness” refers to the amount net radial inward force (i.e., uniform radial inward pressure over the entire abluminal scaffold surface×the abluminal surface area) required to reversibly decrease a scaffold diameter by a certain amount. The slope of the curve from a force-deflection plot will be called the “absolute stiffness” or K. The units are N/mm and the stiffness is expressed for the linearly elastic range of response to the radial force. Thus, for a scaffold deployed to 6.5 mm and having a linear elastic range for radial compression between 6.5 mm and 5.5 mm and a radial stiffness of 20 N/mm, a net inward radial inward force of 10 N is needed to decrease the scaffold diameter from 6.5 mm to 6.0 mm. After the radial force is removed, the scaffold returns to the 6.5 mm diameter.
Alternatively, scaffold radial stiffness may be expressed as a stiffness normalized to the scaffold length, or “length-normalized stiffness” (K-Lnorm). First, the radial deflection is measured for an applied force. Next, for each recorded change in scaffold length, the corresponding applied force is divided by the length of the scaffold. This normalized force (e.g., N/mm) is then used with the displacements to compute a stiffness, rather than the actual force that produced the displacement. The resulting length-normalized stiffness has units of (N/mm per mm). The relationship between K and K-Lnorm for a scaffold with length L is
                              K          ⁢                      -                    ⁢          Lnorm                =                ⁢                  [                                    (                                                F                  ⁢                                                                          ⁢                  2                  ⁢                                      /                                    ⁢                  L                                -                                  F                  ⁢                                                                          ⁢                  1                  ⁢                                      /                                    ⁢                  L                                            )                        *                                          (                                                      D                    ⁢                                                                                  ⁢                    2                                    -                                      D                    ⁢                                                                                  ⁢                    1                                                  )                                            -                1                                              ]                                        =                ⁢                              (                          1              ⁢                              /                            ⁢              L                        )                    *                      [                                          (                                                      F                    ⁢                                                                                  ⁢                    2                                    -                                      F                    ⁢                                                                                  ⁢                    1                                                  )                            *                                                (                                                            D                      ⁢                                                                                          ⁢                      2                                        -                                          D                      ⁢                                                                                          ⁢                      1                                                        )                                                  -                  1                                                      ]                                                  =                ⁢                              (                          1              ⁢                              /                            ⁢              L                        )                    *          K                    
Where D2 is the measured scaffold diameter when uniform radial force F2 is applied and D1 is the measured scaffold diameter when uniform radial force F1 is applied. Hence, K is obtained by multiplying K-Lnorm by the scaffold length L.
Alternatively, scaffold radial stiffness may be normalized both with respect to the scaffold length (L) and the scaffold initial diameter (Do), or “Intrinsic stiffness” (K-norm). The relationships among the three types of radial stiffness areK-norm=(Do)*K-Lnorm=(Do/L)*K 
Similar definitions are adopted for a pinching stiffness, which may be measured by a flat-plate test. Pinching stiffness is discussed in US20110190871. Thus, an absolute, length normalized and intrinsic pinching stiffness, denoted as KP, KP-Lnorm and KP-norm, respectively, for a scaffold of length L and initial height (diameter) Do areKP-norm=(Do)*KP-Lnorm=(Do/L)*KP 
A polymer scaffold can be made from a biodegradable, bioabsorbable, bioresorbable, or bioerodable polymer. The terms biodegradable, bioabsorbable, bioresorbable, biosoluble, or bioerodable refer to the property of a material or stent to degrade, absorb, resorb, or erode away from an implant site. The polymer scaffold is intended to remain in the body for only a limited period of time. In many treatment applications, the presence of a stent in a body may be necessary for a limited period of time until its intended function of, for example, maintaining vascular patency and/or drug delivery is accomplished. Moreover, it has been shown that biodegradable scaffolds allow for improved healing of the anatomical lumen as compared to metal stents, which may lead to a reduced incidence of late stage thrombosis. In these cases, there is a desire to treat a vessel using a polymer scaffold, in particular a bioerodible polymer scaffold, as opposed to a metal stent, so that the prosthesis's presence in the vessel is for a limited duration. However, there are numerous challenges to overcome when developing a polymer scaffold.
Polymer material considered for use as a polymeric scaffold, e.g. poly(L-lactide) (“PLLA”), poly(L-lactide-co-glycolide) (“PLGA”), poly(D-lactide-co-glycolide) or poly(L-lactide-co-D-lactide) (“PLLA-co-PDLA”) with less than 10% D-lactide, and PLLD/PDLA stereo complex, may be described, through comparison with a metallic material used to form a stent, in some of the following ways. A suitable polymer has a low strength to weight ratio, which means more material is needed to provide an equivalent mechanical property to that of a metal. Therefore, struts or fibers must be made thicker and wider to have the required strength for a stent to support lumen walls at a desired radius. The scaffold made from such polymers also tends to be brittle or have limited fracture toughness. The anisotropic and rate-dependent inelastic properties (i.e., strength/stiffness of the material varies depending upon the rate at which the material is deformed) inherent in the material, only compound this complexity in working with a polymer, particularly, bio-absorbable polymer such as PLLA or PLGA.
Scaffolds used to treat coronary vessels experience, for the most part, a primarily radial loading. However, scaffolds intended for peripheral vessels experience a quite different loading, to such an extent that the traditional measure of a stent's fitness for use, i.e., its radial strength/stiffness, is not an accurate measure of whether the scaffold will have sufficient strength to provide mechanical support within the peripheral vessel for the duration needed. This is because a peripheral scaffold is placed in a significantly different environment from a coronary scaffold. The vessel size is larger. And there is much more movement of the vessel, especially when located close to an appendage. As such, a scaffold intended for a peripheral vessel will need to be able to sustain more complex loading, including a combination of axial, bending, torsional and radial loading. See e.g. Bosiers, M. and Schwartz, L., Development of Bioresorbable Scaffolds for the Superficial Femoral Artery , SFA: CONTEMPORARY ENDOVASCULAR MANAGEMENT (‘Interventions in the SFA” section). These and related challenges facing peripherally implanted stents and scaffolds are also discussed in US2011/0190872.
There is a need to develop a prosthesis for treating peripheral blood vessels that can provide mechanical support for the vessel, until this support is no longer needed and then resorb away. There is a further need to develop such a prosthesis that minimizes late lumen loss and stenosis of the vessel, such as within the first month following implantation, thereby providing improved vascular patency.